Abstract

In the last two decades, we have witnessed a remarkable progress in the development of biosensor devices and their application in areas such as environmental monitoring, biotechnology, medical diagnostics, drug screening, food safety, and security, among others. The technology of optical biosensors has reached a high degree of maturity and several commercial products are on the market. But problems of stability, sensitivity, and size have prevented the general use of optical biosensors for real field applications. Integrated photonic biosensors based on silicon technology could solve such drawbacks, offering early diagnostic tools with better sensitivity, specificity, and reliability, which could improve the effectiveness of in-vivo and in-vitro diagnostics. Our last developments in silicon photonic biosensors will be showed, mainly related to the development of portable and highly sensitive integrated photonic sensing platforms.

1. Introduction

The progressive demand for the rapid and precise detection of any type of
substances has speed up the development of a large variety of biosensors. For
most of the applications, it is desirable to have a compact biosensor with high
sensitivity, fast response, and able to perform real-time measurements. These
requirements can be achieved mainly with optical sensors [1] due to the own
nature of the optical measurements that endow a great number of different techniques,
as emission, absorption, fluorescence, refractometry, or polarimetry. Among
them, photonic biosensors based on evanescent wave detection have demonstrated
its outstanding properties, such as an extremely high sensitivity for the
direct measurement of biomolecular interactions, in real time and in label-free
schemes [1].

In the evanescent wave detection, a receptor layer is immobilized onto
the core surface of the waveguide. The exposure of the functionalized surface
to the complementary analyte molecules and the subsequent biochemical
interaction between them induces a local change in the optical properties of
the biological layer. This change is detected via the evanescent field of the
guided light and its amplitude can be correlated to the concentration of the
analyte and to the affinity constant of the interaction, yielding a
quantitative signal of the interaction.

The advantages of optical sensing are significantly improved when this
approach is used within an integrated optics context. Integrated optics technology
allows the integration of passive and active optical components (including
fibres, emitters, detectors, waveguides, and related devices) onto the same
substrate, allowing the flexible development of miniaturized compact sensing
devices, with the additional possibility to fabricate multiple sensors on a
single chip. The integration offers additional advantages such as
miniaturization, robustness, reliability, potential for mass production with
consequent reduction of production costs, low energy consumption, and
simplicity in the alignment of the individual optical elements.

Several technologies are available for the fabrication
of photonic biosensors, but the well-developed silicon technology is one of the
most useful and promising tools. Much work on silicon photonic devices has been
done in the telecommunications field, and some results have been successfully
applied for sensor development. In “siliconized” photonics, the fabrication of
the devices is performed with silicon or silicon-related materials using
microelectronics technologies, with the aim of integrating all the sensing
components in a single chip. In order to “siliconize” photonics,
there are several building blocks for investigation, including light generation
and coupling, selectively guiding and transporting within the waveguides, light
encoding, detection, packaging the devices, and, finally, “smart” electronic control
of all these photonic functions (see Figure 1). Biofunctionalization of the
sensing element is a subject of special research beyond photonics, thus the
element has been highlighted in a separate block on the diagram of Figure 1.

Figure 1: The main research areas in the development of
photonics biosensors based on silicon technology.

In particular, light guiding for sensing applications
can be based on total internal reflection (TIR) planar or rib waveguides [3, 14, 15], hollow waveguides [16–19], antiresonant reflecting optical waveguides (ARROW) [20–22], or slot [4, 23] waveguides. Interestingly,
complex waveguide structures developed for telecommunications with the purpose
of further miniaturization of photonics devices, which show potential sensing
applications, have been reported recently [24].

In this paper, we will discuss the photonic biosensing platforms
based on silicon technology, which are investigating in our laboratory, such as
interferometric Mach-Zehnder (MZI), bimodal waveguides devices, and optomechanical
microcantilevers. The three devices have advantages and disadvantages and
depending on the specific biological application one is more suitable than the
others. The interferometric MZI
sensors show extremely high sensitivity; however, the existing light coupling
methods frequently limit their application to laboratory use only. High thermal
stability is required for the bimodal waveguide interferometers but this device
is simpler in operation than the MZI. The optical microcantilevers also require
a high mechanical and thermal stability for operation but this device is easily
scalable to dozens or hundreds of sensors in the same chip opening the way to
high-throughput screening using label-free biosensors. The main aim of this
article is to describe the use of silicon technology for the implementation of
integrated optical biosensor rather than to perform a sensitivity comparison
between the different devices.

2. Integrated Mach-Zehnder Interferometer (MZI) Sensor

Interferometric biosensors constitute one of the most sensitive integrated-optic
alternatives as compared to other optical biosensor (i.e., plasmonic
biosensors) for label-free detection. In these sensors, the guided light
interacts with the analyte through its evanescent field or, alternatively, the
analyte can propagate in the core of the waveguide if hollow or slot waveguides
are employed. The most common Mach-Zehnder and Young interferometers [3, 25, 26] are composed of an incident
waveguide that is split in two single mode waveguide branches, in which one of
them contains a sensing window.

We
have developed integrated Mach-Zehnder interferometers based on TIR waveguides
(Si/SiO2/Si3N4) of micro/nanodimensions, as it
is shown on Figure 2. For biosensing applications, the optical waveguides of
the MZI must have high surface sensitivity and single mode behavior. For that reason,
we have chosen Si3N4 core layers over a SiO2 substrate . In this
waveguide configuration (and wavelength in the visible range), the single mode
behavior is obtained for core thickness below 300 nm. In order to provide single
mode operation in the lateral direction, rib structures several nanometers high
(below 5 nm) and rib widths of 4 m were fabricated. To achieve single mode behavior, the waveguide could be designed
to be narrower and with a higher rib but this will depend on the tolerances of
the equipment available for the fabrication. With the equipment of our
facilities, we can ensure the reproducibility of the rib height at a nanometric
scale which allows us to make relatively wide waveguides which are more
convenient for experimental work. Figure 2 shows a schematic of the
waveguide configuration. In the MZI, Y-shape divisors with circular arms of mm were designed to direct light in the two branches of the MZI with 3 dB
split ratio. To protect the device from temperature fluctuations, the waveguide
branches are placed very close to each other (100 m,), thus the temperature changes affect both
waveguides simultaneously.

The devices
are fabricated in our clean room facilities. The final device has a length of 3 cm and the sensor area is 1.5 cm long and 50 m
wide. The experimental evaluation of the device was performed in an optical
bench, where polarized light from an He-Ne laser was end-fire coupled to the sensor. The
propagation losses of the waveguides were measured by the Fabry-Perot resonance
technique, and the optical coupling losses were measured by the
cutback method. Propagation losses, in the case of MZI of 200 nm core
thickness, vary between 0.13 and 0.15 dB/cm for TE polarization and between
0.27 and 0.30 dB/cm for TM polarization. Insertion losses are 5.84 dB for TM
polarization and 8.3 dB for TE polarization. The sensitivity of the sensor for both polarizations was analyzed in the
same way as in previous reports [3]. The evaluation was done flowing
solutions of water and ethanol of varying concentration (refractive index steps
of 10−3) and measuring the output signal in real time. Taking into account
the signal-to-noise ratio of our system, the lowest detection limit in the
variation of the refractive index for the TM polarization was
found to be .

The relatively simple design, high integration level,
well developed read out techniques, and the high sensitivity make these devices
very attractive for bio/chemical applications. The MZI device has been used for
the direct detection of DNA hybridization and for the detection of single mutations
at the BRCA-1 gene, involved in breast cancer development, without target labelling
[27]. The oligonucleotide probe is immobilized by covalent attachment to the
sensor surface through silanization procedures. A silane (3-mercaptopropyltrimethoxysilane)
with a thiol group at the free end was employed for the chemical modification
of the surface. A thiol-derivatized oligonucleotides (28 mer) used as receptors
can bind to the silanized Si3N4 surface through a
disulphide bond. The DNA probe has also a 15-tiamine tail which is employed as a vertical
spacer chain to increase the accessibility to the complementary DNA to the
sensor surface.

After DNA immobilization,
complementary oligonucleotides (58 mer) were flowed in the sensor for
hybridization experiments. The hybridization was performed for different DNA
target concentrations from 1 pM to 1 M. Regeneration after each hybridization was
achieved flowing deionized water and HCl 3.2 mM. Figure 3 shows examples of the
real-time detection of DNA hybridization for several concentrations. The calibration curve as a
function of the DNA concentration can be seen in Figure 4. In these measurements, a 10 pM-concentration of complementary nonlabelled DNA in buffer solution was the
lowest hybridization limit achieved, which means an average DNA growth layer of nm, corresponding to an
estimation of DNA molecules/cm2 hybridized at the sensor area of the MZI. In contrast, noncomplementary oligonucleotides
did not show any significant signal.

More importantly, we have
detected the hybridization of 100 nM DNA target with two mismatching bases
corresponding to a mutation of the BRCA-1 gene (data not shown). These results
place the Mach-Zehnder interferometer as one of the most sensitive optical
biosensor for label-free mismatch and DNA hybridisation detection.

3. Bimodal Waveguide Sensor

This sensor is comprised by a single straight bimodal
waveguide (BiMW), which supports the zero- and first-order transversal modes
(see Figure 5). These modes propagate with different velocities depending,
among other factors, on the refractive index of the cladding layer. The
interference pattern formed at the exit of the waveguide changes if the
refractive index varies. The pattern is projected on a two-sectional
photodetector (TSP), then the intensity maximum moves between the lower and the
upper sections of the photodetector. The signals generated by the photodetector
sections are recalculated into a parameter ,
according to the expression where are the signals generated by the upper and the lower sections of the
photodetector, respectively.

Figure 5: Schematic view of the bimodal waveguide sensor
device. The modes are propagating with different velocities. The interference
pattern, created at the exit and projected onto a two-sectional photodetector
(TSP), varies as a function of the refractive index of the cladding layer.

The difference between can reach 17 dB,
which means a variation of from
0 to 0.96. These values were calculated assuming a silicon nitride waveguide
with thickness of 400 nm operated at 633 nm. The total output signal, which can
be represented by the denominator in the right part in (1), is proportional to
the light power coupled into the waveguide, except for minor changes due to
reflection at the output facet, which slightly depends on the intensity
distribution at the exit, according to the simulations. Using the parameter , the
ambiguities due to coupling efficiency variations can be significantly reduced.
However, as the monitoring of light power coupled into the waveguide is still
desirable, a part of the incoupled light can be tapped off and measured with a
conventional photodetector. The sensitivity of the device is given by
where is the length of the sensing window, is the wavelength, are the effective refractive index of the zero-
and first-order modes, respectively, is the phase shift between
the modes, and is the refractive index of the cladding layer.

For the
experiments, we employed a 3 m-wide, 400 nm thick Si3N4 waveguide deposited on a silicon dioxide buffer layer, and the length of the
sensing area was 3 mm.
The waveguide was excited by direct focusing of light from an He-Ne laser (633 nm, 10 mW). A slight misalignment of the objective with respect to the
waveguide in the vertical direction allows for excitation of both modes
simultaneously. Light was collected by an objective lens, and the image of the
waveguide facet was projected on the photodetector (TSP).

The detection of the refractive index changes was
performed by injecting varying concentrations of water/glycerin solutions into
a microfluidic channel formed over the sensing area. An example of the changes
of the interferogram due to refractive index variation is shown in Figure 6.
The sensitivity, defined as the relative change in the output signal per RIU
change, can reach a value of better than per RIU on a 1 cm-long
waveguide with thickness less than 400 nm. A sensitivity of per RIU was demonstrated experimentally on a silicon nitride waveguide with
thickness of 400 nm. The obtained sensitivity was limited mostly by the
thickness of the waveguide and by the coupling technique which allowed for only
25% modulation of the output signal. Biosensing experiments with this device
are in progress.

Figure 6: Response of the BiMW
sensor to the injection of glycerin solutions with concentrations of 2.2% and
3.3% (v).

4. Optical Microcantilever Biosensor

The development of microprobes
for atomic force microscopy (AFM) was an important milestone for the establishment
of efficient technological approaches to MEMS sensors. The principle of operation
of the microcantilever sensor is based on the bending induced in the cantilever
when a biomolecular interaction takes place in one of its surfaces. In this
way, microcantilevers translate the molecular reaction into a nanomechanical
motion, which is commonly detected using optical or piezoresistive readout [28].

In order to achieve highly
integrated microsystem with microcantilever transducers, we have recently
introduced a new type of readout technique. The combination of
photonics and mechanics has been demonstrated in an optical waveguide
cantilever sensor. The sensor can work in static or dynamic
modes, either by monitoring the deflection or the changes in the resonance
frequency of the cantilever. The principle of operation is based on the
sensitivity of energy transfer between two butt-coupled waveguides to their
misalignment with respect to each other as it is represented in Figure 7. The
advantage of the device is that the transducer is integrated with the receptor in
the same chip and the external photodetector is only used for optical power
readout. No preliminary alignment or adjustment is needed, except for light
coupling into the chip.

Figure 7: Sketch of the sensor based on
optical waveguide microcantilevers.

We have fabricated arrays of 20 optical microcantilevers. Each of
them is 200 m long, 40 m wide, and 500 nm thick with a spring constant
of 0.050 N/m. Fabrication of the sensor is done using standard microelectronic
technology. The cantilevers are made of thermal silicon dioxide, transparent in
visible range. Input and output waveguides are made of silicon nitride and are 140 nm-thick and 40 m-wide. The cantilever has low stress gradient
and is practically flat, the misalignment between the output waveguide and the
cantilever free end is around 1 m. Some photographs of
the fabricated devices can be seen in Figure 8. Coupling of the light in the
cantilever is achieved through the evanescent field of the input waveguide.

Figure 8: Photographs of the optical
cantilevers and the light coupling inside them.

In order to characterize
the sensor, an experimental setup was used to measure the amplitude of
modulation of the output signal induced by the vibration of the cantilever at
the resonance frequency. Light from He-Ne laser (632.8 nm, 7.5 mW) was coupled
into the chip using direct focusing with an objective lens (40x, NA 0.65) and
was collected upon exiting by another objective (40x, NA 0.65) before being
directed to a silicon photodetector connected to an oscilloscope and an
acquisition system for spectrum analysis through a low-noise amplifier with
bandwidth 5 to 45 kHz, at FWHM. Light from the same laser source after
splitting was focused by a lens with a focal distance of 75 mm on the cantilever near
its free end. The reflected beam was projected on to a two-sectional position
sensitive photodetector to monitor the displacement of the cantilever. Clear
resonance behavior near 13 kHz with a -factor of 12 was observed [11]. The
change in the output voltage per unit cantilever displacement was calculated to
be 15 V/nm, thus nanometre resolution of the system
was demonstrated.

Taking account in this configuration
that the
minimum detectable deflection is limited by the shot noise of the
photodetector, the Johnson noise of the load resistor, the noise in the
acquisition system, the cantilever vibration due to the thermal noise, and the
noise produced by the laser source, we have estimated that, for a 1 m wide gap, the cantilever displacement can be detected with
a resolution of 18 [11], showing similar performances for
biosensing than the standard microcantilevers.

One of the advantages of this device is that real-time
parallel monitoring of several channels can be done simultaneously, opening the
way for multisensing. The cantilever can be metallized with gold (and then it
is possible to use the well-known thiol chemistry for bioreceptors
immobilization) and its initial displacement can be adjusted by varying the
power of light coupled inside [29]. This new device has
shown good performances for biosensing and offers an interesting approach for
further integration in lab-on-a-chip microsystems. The integration with the light source and the
biofunctionalization of the device is a subject of our current research.

5. Integration in “Lab-on-a-Chip” Microsystems

For the development of a complete photonic
lab-on-a-chip microsystem device, several units must be incorporated on the
same platform: (i) the micro/nanodevices, (ii) the flow cells and the flow
delivery system, (iii) for interferometric sensors, a phase modulation system
to convert the periodic output signal in direct phase measurements, (iv)
integration of the light sources and the photodetectors, and (v) CMOS
processing electronics. For achieving this goal, our first step has been the
development of a novel low temperature (100°C) CMOS compatible microfluidic technology
to create 3D embedded interconnected microfluidic channels between different
substrates. The microfluidic channels have a height from 40 to 60 m and a width
between 100 to 250 m. More details can
be found in [30].

6. Conclusions

We have
presented the development of different integrated optical biosensor platforms
based on silicon technologies: a Mach-Zehnder integrated interferometer, a
bimodal waveguide sensor, and a waveguided microcantilever device. The feasibility
of the different platforms for biosensing has been proved. In the case of the
MZI device, we have achieved a lowest limit of detection of 10 pM for the DNA hybridization
of the BRCA-1 gene, involved in breast cancer development. These results place
the Mach-Zehnder interferometer as one of the most sensitive optical biosensor
for label-free DNA detection.

All the
described sensing configurations look ahead the possibility of the integration of
optic, fluidic, and electrical functions on one platform to obtain lab-on-a-chip
microsystems. These results open the way for further development of
portable and multianalyte sensors for the detection of several biological molecules
of interest in situ and in real-time.

Acknowledgments

Authors would like to thank the financial support
of the Spanish Ministry of Education and Science (project TEC2005-13604) and M.
Botín Foundation.